Polyester (DACRON® or polyethylene terephthalate) fibers were first characterized in 1941 and have become the most widely produced synthetic fiber in the world. They are most familiarly known by the DuPont commercial name DACRON®. The polymer is synthesized by a condensation reaction of derivatives of ethylene glycol and terephthalic acid, resulting in molecules that contain 80 to 100 repeat units. These molecules are then extruded through a plurality of holes (a spinneret) to produce multi-filament fibrous filaments. Such DACRON® fibers are further processed into various structures such as warp-knit, weft-knit, and woven fabrics that have excellent resiliency as well as resistance to a wide range of chemical and biological challenges.
DACRON® is utilized in items ranging from clothing to medical implants. DACRON® yarn was first sewn into a tubular form and utilized as a large-diameter vascular graft in the mid-1950s. Since that point, DACRON® has been incorporated into both large and medium bore vascular grafts in knitted and woven form. These grafts have shown excellent long-term biodurability, handling characteristics and capsular tissue incorporation.
Polyester is known as a relatively inert fiber. It is hydrophobic, both in bulk and in its surface properties. At normal temperatures it has low uptake of moisture, dyes and other chemicals. In normal textile use, it tends to suffer associated disadvantages: it generates static electricity, it does not readily shed oily soils, and it does not wet enough to encourage the wicking of water. For applications where repellency is required, however, it is insufficiently hydrophobic, and repellent finishes are applied. Like any fiber, softness is advantageous, and chemical softeners are applied.
Modifications to overcome these deficiencies typically rely on the surface deposition of polymeric textile finishes. These include silicones, vinyl and acrylic polymers, and fluorochemicals. Other finishes are based on an ester interchange reaction that fixes a hydrophilic moiety (typically a short chain polyethylene oxide). Many of these finishes suffer a lack of durability to laundering and dry cleaning, since (other than those bonded via ester interchange) they are not covalently bonded to the polyester surface.
Polyester is employed for various medical devices such as prosthetic vascular grafts, prosthetic heart valve sewing cuffs, left ventricular assist devices, artificial organs, wound patches and wound dressings. Polyester is a biodurable material due to the relatively inert properties of the polymer and can persist for greater than 10 years when implanted without significant deleterious effects to the specific device. However, this material, similar to other biomaterials, is prone to 3 major complications when implanted: 1) thrombosis (clot formation), 2) infection and 3) lack of cell appropriate healing. These adverse properties occur as a result of the bulk properties of the polymer. Additionally, the rigidity of the polymer limits surface modification in order to incorporate various moieties such as anti-thrombolytic agents (e.g., anti-thrombin), thrombolytic agents, growth-promoting factors, growth-inhibiting factors, and antimicrobial/antifungal agents.
A complication of all implantable biomaterials is incompatibility between blood and the biomaterial surface. The initial interaction of blood and the foreign surface results in an array of activation or biologic responses: platelet activation and adhesion, activation of the intrinsic pathway of the coagulation cascade resulting in formation of active thrombin, leukocyte activation and the release of complement and kallikrein. If unregulated, these responses lead to surface thrombus formation with subsequent failure of the implanted biomaterial.
Numerous attempts have been made to create a more biocompatible surface by establishing a new biologic lining on the luminal surface that would “passivate” this acute initial reaction. These efforts have ranged from non-specifically binding albumin to the surface followed by heat denaturation to non-specifically crosslinking albumin, gelatin and collagen. Covalent or ionic binding of the anticoagulant heparin alone, in conjunction with other biologic compounds, or with spacer moieties as well as covalent linkage of thrombomodulin have also been performed. Other studies have focused on modifying the composition of the biomaterial by either increasing hydrophilicity via incorporation of polyethylene oxide groups or by creating an ionically charged surface.
Each of these methodologies has had limited success in creating a durable, biologically-active surface. There are several limitations associated with these surface modifications: 1) thrombin is not directly inhibited therefore fibrinogen amounts remain constant on the material surface permitting platelet adhesion, 2) heparin-coated biomaterials may be subject to heparitinases limiting long-term use of these materials, 3) non-specifically bound compounds are desorbed from the surface which is under shear stresses thereby re-exposing the thrombogenic biomaterial surface, 4) rapid release of non-specifically bound compounds may create an undesired systemic effect and 5) charge-based polymers may be covered by other blood proteins such that anticoagulant effects are masked.
Endothelial cells play a pivotal role in mediating blood interaction with the arterial wall. These cells maintain hemostasis and also synthesize growth mediators that block abnormal smooth muscle cell proliferation. Ideally, prosthetic grafts should promote endothelial cell adherence and growth on the luminal surface while permitting direct host tissue incorporation at the capsular surface. This type of cellular incorporation does not occur in actuality, thereby predisposing these grafts to infection, thrombosis, perigraft seromas and delayed graft failure. Thus, failure of appropriate cell type growth and development to these biomaterials significantly limits their expanded use.
To avert these complications and mimic the non-thrombogenic in vivo endothelial cell blood vessel lining, cell adhesion to prosthetic grafts using endothelial cell seeding techniques have been extensively employed. Adhesive proteins such as fibronectin, fibrinogen, vitronectin and collagen have served well in graft seeding protocols. The cell attachment properties of these matrices can also be duplicated by short peptide sequences such as RGD (Arg-Gly-Asp). These adhesive proteins, however, have several drawbacks: 1) bacterial pathogens recognize and bind to these sequences, 2) non-endothelial cell lines also bind to these sequences, 3) patients requiring a seeded vascular graft have few donor endothelial cells, therefore cells must be grown in culture and 4) endothelial cell loss to shear forces remains a significant obstacle.
Modification of the surface has also been employed to modify host response to the foreign body, serving as an approach for improving endothelial cell adherence to DACRON®. Endothelial cells after seeding have been shown to attach and grow on a variety of protein substrates coated onto vascular graft materials. Bioactive oligopeptides and cell growth factors have been immobilized onto various polymers and demonstrated to affect cell adherence and growth. Additional studies have described the incorporation of growth factors into a degradable protein mesh, resulting in the formation of capillaries into the graft wall. Utilizing these techniques to incorporate growth factors, however, does have major limitations: 1) growth factor is rapidly released from the matrix, 2) matrix degradation re-exposes the thrombogenic surface, thus endothelialization is not uniform and 3) release of non-endothelial specific growth factor is not confined to the biomaterial matrix, thereby exposing the “normal” distal artery to the growth factor.
There have been several studies assessing the effects of amine interaction with polyester. Zahn et al. (Polymer 3:429, 1962), as well as Farrow et al. (Polymer 3:17, 1962) assessed the lysis of polyester in an attempt to breakdown excess material in the textile industry into smaller components, without regard to maintaining the integrity of the polymer structure. In 1982, Ellison et. al. examined the effects of a monofunctional amine versus alkali hydrolysis on polyester. These studies, which again were performed under harsh conditions, demonstrated that alkaline hydrolysis has a more substantial effect on fiber weight without extensive strength loss. In contrast, aminolysis had less effect on fiber weight but a greater effect on fiber strength, indicative of a permanent reaction within the polymer structure. In 1968, Avny and Rebenfeld demonstrated that multi-functional amine compounds could be reacted within the polymer structure (three or more amine groups) while presenting minimal loss in strength (Applied Polymer Science 32:4009, 1986). There remains a need, however, for a polyester material that provides functional moieties for attachment of commercial finishes or biologically-active agents, while retaining material strength.